U.S. patent application number 14/805626 was filed with the patent office on 2016-01-21 for implantable pulse generator that generates spinal cord stimulation signals for a human body.
The applicant listed for this patent is GLOBUS MEDICAL, INC.. Invention is credited to Raghavendra Angara, Christopher Biele, Saif Khalil.
Application Number | 20160015981 14/805626 |
Document ID | / |
Family ID | 55073701 |
Filed Date | 2016-01-21 |
United States Patent
Application |
20160015981 |
Kind Code |
A1 |
Biele; Christopher ; et
al. |
January 21, 2016 |
IMPLANTABLE PULSE GENERATOR THAT GENERATES SPINAL CORD STIMULATION
SIGNALS FOR A HUMAN BODY
Abstract
An implantable pulse generator (IPG) that generates spinal cord
stimulation signals for a human body includes a timing generator
and high frequency generator. The timing generator generates timing
signals that represent stimulation signals for multiple channels.
The high frequency generator determines whether to modulate the
timing signals and modulates them at a burst frequency according to
stored burst parameters if the decision is yes. The high frequency
generator can also independently control the pulse frequency of
each channel according to the stored parameters. As such, the IPG
provides the ability to generate both the low frequency and high
frequency stimulation signals at different frequencies in different
channels according to user programming in order to provide maximum
flexibility in treatment.
Inventors: |
Biele; Christopher; (King of
Prussia, PA) ; Angara; Raghavendra; (Norristown,
PA) ; Khalil; Saif; (Wayne, PA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
GLOBUS MEDICAL, INC. |
AUDUBON |
PA |
US |
|
|
Family ID: |
55073701 |
Appl. No.: |
14/805626 |
Filed: |
July 22, 2015 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
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14213186 |
Mar 14, 2014 |
|
|
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14805626 |
|
|
|
|
61792654 |
Mar 15, 2013 |
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Current U.S.
Class: |
607/46 |
Current CPC
Class: |
A61N 1/025 20130101;
A61N 1/0553 20130101; A61N 1/36185 20130101; A61N 1/37229 20130101;
A61N 1/37235 20130101; A61N 1/3752 20130101; A61N 1/3787 20130101;
A61N 1/36178 20130101; A61N 1/36071 20130101; A61N 1/36125
20130101 |
International
Class: |
A61N 1/36 20060101
A61N001/36; A61N 1/372 20060101 A61N001/372 |
Claims
1. An implantable pulse generator (IPG) that generates spinal cord
stimulation signals for a human body, comprising: control registers
storing stimulation signal parameters for a plurality of channels
with each channel capable of being associated with at least two
electrodes among a plurality of electrodes, each channel
representing a stimulation signal pattern of the associated
electrodes, the signal parameters including burst parameters that
specify the frequency of the stimulation signal for each channel; a
timing generator adapted to generate timing signals representing
stimulation signals for the plurality of channels according to the
stored signal parameters; a high frequency generator adapted to
determine whether to modulate the timing signals received from the
timing generator and modulate the received timing signals at a
burst frequency according to the stored burst parameters so as to
enable independent frequency control of the stimulation signals for
each channel.
2. The IPG of claim 1, further comprising an electrode driver for
each of the plurality of electrodes, and adapted to adjust the
amplitude of the timing signals from the timing generator according
to the stored signal parameters so as to enable independent
amplitude control of the stimulation signals for each channel.
3. The IPG of claim 1, wherein when the high frequency generator
has determined not to modulate the received timing signals, the
high frequency generator outputs the received timing signals whose
frequency is unaltered.
4. The IPG of claim 1, wherein the high frequency generator
includes a burst multiplexer that: receives the burst parameters
stored in the control registers; selects the burst parameters
associated with an active channel; and outputs the selected burst
parameters.
5. The IPG of claim 4, wherein the high frequency generator
includes a burst generator that receives the selected burst
parameters from the burst multiplexer and modulates the received
timing signals according to the received burst parameters.
6. The IPG of claim 1, further comprising: a processor having a
memory; a treatment control module stored in the memory and
executable by the processor, the treatment control module adapted
to transmit the stimulation signal parameters to the control
registers and to place the processor in a standby mode, wherein the
timing generator is adapted to generate the timing signals
according to the stored signal parameters without intervention from
the processor while the processor is in the standby mode.
7. The IPG of claim 6, further comprising a transceiver module
adapted to wirelessly communicate with a remote control and to
transmit an interrupt signal to the processor to wake up the
processor from the standby mode.
8. The IPG of claim 1, further comprising an arbitrator that
continuously receives the generated timing signals and selects one
channel among the plurality of channels as an active treatment
channel.
9. The IPG of claim 8, wherein when the one channel is being
selected as the active treatment channel, the arbitrator is adapted
to suppress all other channels.
10. The IPG of claim 8, wherein when the generated timing signals
indicate that two or more channels are about to become active at
the same time when no active channel has been selected, the
arbitrator is adapted to select the one channel as the active
treatment channel based on a predetermined priority of the two or
more channels.
11. A method of generating spinal cord stimulation signals for a
human body, comprising: storing in control registers stimulation
signal parameters for a plurality of channels with each channel
capable of being associated with at least two electrodes among a
plurality of electrodes, each channel representing a stimulation
signal pattern of the associated electrodes, the signal parameters
including burst parameters that specify the frequency of the
stimulation signal for each channel; generating timing signals
representing stimulation signals for the plurality of channels
according to the stored signal parameters; determining whether to
modulate the timing signals received from the timing generator; and
if it is determined that the timing signals are to be modulated,
modulating the received timing signals at a burst frequency
according to the stored burst parameters so as to enable
independent frequency control of the stimulation signals for each
channel.
12. The method of claim 11, further comprising: for each of the
plurality of electrodes, adjusting the amplitude of the timing
signals from the timing generator according to the stored signal
parameters so as to enable independent amplitude control of the
stimulation signals for each channel.
13. The method of claim 11, wherein if it is determined that the
timing signals are not to be modulated, outputting the received
timing signals whose frequency is unaltered.
14. The method of claim 11, further comprising: selecting the burst
parameters associated with an active channel among the plurality of
channels, wherein the step of modulating includes modulating the
received timing signals according to the selected burst
parameters.
15. The method of claim 11, further comprising: transmitting, by a
processor, the stimulation signal parameters to the control
registers; and placing the processor in a standby mode, wherein the
timing signals are generated without intervention from the
processor while the processor is in the standby mode.
16. The method of claim 15, further comprising receiving by the
processor an interrupt signal from a transceiver module to wake up
the processor from the standby mode.
17. The method of claim 11, further comprising continuously
receiving the generated timing signals and selecting one channel
among the plurality of channels as an active treatment channel.
18. The method of claim 17, wherein when the one channel is being
selected as the active treatment channel, further comprising
suppressing all other channels.
19. The method of claim 17, wherein when the generated timing
signals indicate that two or more channels are about to become
active at the same time when no active channel has been selected,
further comprising selecting the one channel as the active
treatment channel based on a predetermined priority of the two or
more channels.
20. The method of claim 11, wherein the signal parameters include
an identification of all electrodes that belong to each channel and
in which each electrode can be associated with two or more
channels.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] The present application is a continuation-in-part of U.S.
patent application Ser. No. 14/213,186, filed Mar. 14, 2014, which
claims priority to U.S. Provisional Application Ser. No.
61/792,654, filed Mar. 15, 2013, and entitled "SPINAL CORD
STIMULATOR SYSTEM," all of which are herein incorporated by
reference in their entirety.
TECHNICAL FIELD
[0002] This disclosure relates to stimulators using electrical
pulses in a medical context, and more particularly, applying
electrical stimulation signals to the spinal cord to control
pain.
BACKGROUND
[0003] A Spinal Cord Stimulator (SCS) is used to exert pulsed
electrical signals to the spinal cord to control chronic pain.
Spinal cord stimulation, in its simplest form, comprises
stimulating electrodes implanted in the epidural space, an
implantable pulse generator implanted in the lower abdominal area
or gluteal region, conducting wires connecting the electrodes to
the electrical pulse generator, an electrical pulse generator
remote control, and an electrical pulse generator charger. Spinal
cord stimulation has notable analgesic properties and, at the
present, is used mostly in the treatment of failed back surgery
syndrome, complex regional pain syndrome and refractory pain due to
ischemia.
[0004] Electrotherapy of pain by neurostimulation began shortly
after Melzack and Wall proposed the gate control theory in 1965.
This theory proposed that nerves carrying painful peripheral
stimuli and nerves carrying touch and vibratory sensation both
terminate in the dorsal horn (the gate) of the spinal cord. It was
hypothesized that input to the dorsal horn of the spinal cord could
be manipulated to "close the gate" to the nerves. As an application
of the gate control theory, Shealy et al. implanted the first
spinal cord stimulator device directly on the dorsal column for the
treatment of chronic pain in 1971.
[0005] Spinal cord stimulation does not eliminate pain. The
electrical pulses from the stimulator override the pain messages so
that the patient does not feel the pain intensely. In essence, the
stimulator masks the pain. A trial implantation is performed before
implanting the permanent stimulator. The physician first implants a
trial stimulator through the skin (percutaneously) to perform
stimulations as a trial run. Because a percutaneous trial
stimulator tends to move from its original location, it is
considered temporary. If the trial is successful, the physician can
then implant a permanent stimulator. The permanent stimulator is
implanted under the skin of the abdomen with the leads inserted
under the skin and subcutaneously fed to and inserted into the
spinal canal. This placement of the stimulator in the abdomen is a
more stable, effective location. The leads, which consist of an
array of electrodes, can be percutaneous type or paddle type.
Percutaneous electrodes are easier to insert in comparison with
paddle type, which are inserted via incision over spinal cord and
laminectomy.
[0006] There are a number of problems that exist in currently
available implantable pulse generators that limit the full benefits
of dorsal column imulation from an effectiveness and patient user
friendly perspective.
[0007] One problem is that the circuits in the current generators
consume too much power. This requires frequent recharging, making
it very inconvenient for patients. Another problem is that the
current generators are limited in concurrently generating different
stimulation patterns to treat different parts of the body
simultaneously. Accordingly, when patients have varying degrees of
pain in different parts of the body, it is difficult, if not
impossible, to affectively treat all area of pain.
[0008] Therefore, it would be desirable to provide a system and
method for generating stimulation patterns which resolve the
problems discussed above.
SUMMARY OF THE DISCLOSURE
[0009] According to one aspect of the present invention, there is
provided an implantable pulse generator (IPG) that generates spinal
cord stimulation signals for a human body has a programmable signal
generator that can generate the signals based on stored signal
parameters without any intervention from a processor that controls
the overall operation of the IPG. While the signal generator is
generating the signals the processor can be in a standby mode to
substantially save battery power.
[0010] According to another aspect of the present invention, there
is provided an implantable pulse generator (IPG) that generates
spinal cord stimulation signals for a human body has control
registers that store stimulation signal parameters for stimulation
channels with each channel capable of being associated with at
least two electrodes and representing a particular stimulation
signal pattern for the associated electrodes. An arbitrator
continuously receives timing signals representing the stimulation
signal patterns and selects one channel among the many channels as
an active treatment channel in order to avoid two channels from
being activated at the same time. The arbitrator provides
flexibility in programming different pulse parameters for multiple
stimulation channels without the possibility of overloading the
power supply that generates the stimulation signal patterns.
[0011] According to another aspect of the present invention, there
is provided an implantable pulse generator (IPG) that generates
spinal cord stimulation signals for a human body, which includes a
timing generator and high frequency generator. The timing generator
generates timing signals that represent stimulation signals for
multiple channels. The high frequency generator determines whether
to modulate the timing signals and modulates them at a burst
frequency according to stored burst parameters if the decision is
yes. As such, the IPG provides the ability to generate both the low
frequency and high frequency stimulation signals in different
channels according to user programming.
[0012] According to another aspect of the present invention, there
is provided an implantable pulse generator (IPG) that generates
spinal cord stimulation signals for a human body, which includes a
timing generator and high frequency generator. The timing generator
generates timing signals that represent stimulation signals for
multiple channels. The high frequency generator determines whether
to modulate the timing signals and modulates them at a burst
frequency according to stored burst parameters if the decision is
yes. The high frequency generator can also independently control
the pulse frequency of each channel according to the stored
parameters. As such, the IPG provides the ability to generate both
the low frequency and high frequency stimulation signals at
different frequencies in different channels according to user
programming in order to provide maximum flexibility in
treatment.
[0013] According to another aspect of the present invention, there
is provided an implantable pulse generator (IPG) that generates
spinal cord stimulation signals for a human body, which includes an
electrode driver for each electrode, which adjusts the amplitude of
the timing signals and output an output current corresponding to
the adjusted signals for transmission to the associated electrode
so as to enable independent amplitude control of the stimulation
signals for each stimulation pattern channel.
BRIEF DESCRIPTION OF THE DRAWINGS
[0014] FIG. 1 depicts various components that can be included in a
spinal cord stimulation system, according to an embodiment, during
trial and permanent implantation.
[0015] FIG. 2 depicts an exploded view of an implantable pulse
generator (IPG) assembly, according to an embodiment.
[0016] FIG. 3 depicts a feedthrough assembly of an implantable
pulse generator (IPG) assembly, according to an embodiment.
[0017] FIG. 4 depicts a lead contact system of an implantable pulse
generator (IPG) assembly, according to an embodiment.
[0018] FIG. 5 depicts a lead contact assembly of an implantable
pulse generator (IPG) assembly, according to an embodiment.
[0019] FIG. 6 depicts a head unit assembly of an implantable pulse
generator (IPG) assembly, according to an embodiment.
[0020] FIG. 7 depicts an RF antenna of an implantable pulse
generator (IPG) assembly, according to an embodiment.
[0021] FIG. 8 depicts a percutaneous lead, according to an
embodiment.
[0022] FIG. 9 depicts a paddle lead, according to an
embodiment.
[0023] FIG. 10 depicts a lead extension, according to an
embodiment.
[0024] FIG. 11 depicts a lead splitter, according to an
embodiment.
[0025] FIG. 12 depicts a sleeve anchor, according to an
embodiment.
[0026] FIG. 13 depicts a mechanical locking anchor, according to an
embodiment.
[0027] FIG. 14 illustrates communication via a wireless dongle with
a tablet/clinician programmer and smartphone/mobile/patient
programmer during trial and/or permanent implantation, according to
an embodiment.
[0028] FIG. 15 depicts a Tuohy needle, according to an
embodiment.
[0029] FIG. 16 depicts a stylet, according to an embodiment.
[0030] FIG. 17 depicts a passing elevator, according to an
embodiment.
[0031] FIG. 18 depicts a tunneling tool, according to an
embodiment.
[0032] FIG. 19 depicts a torque wrench, according to an
embodiment.
[0033] FIG. 20 is a function block diagram of some components in an
implantable pulse generator according to an embodiment.
[0034] FIG. 21 is a functional block diagram of the signal
generator of FIG. 20.
[0035] FIG. 22 is a functional block diagram of the high frequency
generator of FIG. 21.
[0036] FIG. 23 is a functional block diagram of the electrode
driver of FIG. 21.
[0037] FIG. 24 is a functional illustration of two of the current
drivers of FIG. 23.
[0038] FIG. 24 shows exemplary electrode waveforms for an active
channel according to an embodiment of the present invention.
[0039] FIG. 25 illustrates a grouping of electrodes for different
channels according to an embodiment of the present invention.
[0040] FIG. 26 shows exemplary electrode waveforms illustrating an
asymmetrical pulse pattern between a positive and negative pulse
according to an embodiment of the present invention.
[0041] FIG. 27 shows exemplary electrode waveforms for two
arbitrated stimulation channels according to an embodiment of the
present invention.
DETAILED DESCRIPTION
Implantable Pulse Generator (IPG)
[0042] FIG. 1 illustrates various components that can be included
in a SCS system for the trial and the permanent installation
periods. The spinal cord stimulator (SCS) 100 is an implantable
device used to deliver electrical pulse therapy to the spinal cord
in order to treat chronic pain. The implantable components of the
system consist of an Implantable Pulse Generator (IPG) 102 and a
multitude of stimulation electrodes 130. The IPG 102 is implanted
subcutaneously, no more than 30 mm deep in an area that is
comfortable for the patient while the stimulation electrodes 130
are implanted directly in the epidural space. The electrodes 130
are wired to the IPG 102 via leads 140, 141 which keep the
stimulation pulses isolated from each other in order to deliver the
correct therapy to each individual electrode 130.
[0043] The therapy delivered consists of electrical pulses with
controlled current amplitude ranging from +12.7 to -12.7 mA
(current range 0-25.4 mA). These pulses can be programmed in both
length and frequency from 100 to 20000 and 0.5 Hz to 1200 Hz. At
any given moment, the sum of the currents sourced from the anodic
electrodes 130 must equal the sum of the currents sunk by the
cathodic electrodes 130. In addition, each individual pulse is
bi-phasic, meaning that once the initial pulse finishes another
pulse of opposite amplitude is generated after a set holdoff
period. The electrodes 130 may be grouped into stimulation sets in
order to deliver the pulses over a wider area or to target specific
areas, but the sum of the currents being sourced at any one given
time may not exceed 20 mA. A user can also program different
stimulation sets (up to eight) with different parameters in order
to target different areas with different therapies.
[0044] FIG. 2 depicts an exploded view of an IPG 102. The IPG 102
consists of two major active components 104, 106, a battery 108,
antenna 110, some support circuitry, and a multitude of output
capacitors 112. The first of the major active components is the
microcontroller 104 transceiver 104. It is responsible for
receiving, decoding, and execution both commands and requests from
the external remote. If necessary it passes these commands or
requests onto the second major component, the ASIC 106. The ASIC
106 receives the digital data from the microcontroller 104 and
performs the entire signal processing to generate the signals
necessary for stimulation. These signals are then passed onto the
stimulation electrodes 130 in the epidural space.
[0045] The ASIC 106 is comprised of a digital section and an analog
section. The digital section is divided into multiple sections
including; Timing Generators, Arbitration Control, Pulse Burst
Conditioner, and Electrode Logic. The analog section receives the
incoming pulses from the digital section and amplifies them in
order to deliver the correct therapy. There are also a multitude of
digital register memory elements that each section utilizes, both
digital and analog.
[0046] The digital elements in the ASIC 106 are all made up of
standard subsets of digital logic including logic gates, timers,
counters, registers, comparators, flip-flips, and decoders. These
elements are ideal for processing the stimulation pulses as all of
them can function extremely fast--orders of magnitudes faster than
the required pulse width. The elements all function at one single
voltage, usually 5.0, 3.3, 2.5, or 1.8 volts.
[0047] The timing generators are the base of each of the
stimulation sets. It generates the actual rising and falling edge
triggers for each phase of the bi-phasic pulse. It accomplishes
this by taking the incoming clock that is fed from the
microcontroller 104 and feeding it into a counter. For the purpose
of this discussion, assume the counter simply counts these rising
clock edges infinitely. The output of the counter is fed into six
different comparators. The comparators other input is connected to
specific registers that are programmed by the microcontroller 104.
When the count equals the value stored in the register, the
comparator asserts a positive signal.
[0048] The first comparator is connected to the SET signal of a SR
flip flop. The SR flip flop stays positive until the RESET signal
is asserted, which the second comparator is connected to. The
output of the SR flip flop is the first phase of the bi-phasic
pulse. Its rising & falling edges are values stored in the
registers and programmed by the microcontroller 104. The third and
fourth comparators & registers work in exactly the same way to
produce the second phase of the bi-phasic pulse using the second SR
flip flop.
[0049] The fifth comparator is connected the RESET of the final
SR-Flip flop in the timing generator. This flip flop is SET by the
first comparator, which is the rising edge of the first pulse. The
RESET is then triggered by the value the microprocessor programmed
into the register connected to the comparator. This allows for a
`holdoff` period after the falling edge of the second pulse. The
output of this third SR flip flop can be thought of as an envelope
of the biphasic pulses indicating when this particular timing
generator is active.
[0050] The final comparator of the system is once again connected
to a register that stores the frequency values from the
microprocessor. Essentially when the count reaches this value it
triggers the comparator which is fed back to the counter to reset
it to zero and beginning the entire pulse generation cycle again.
The ASIC 106 may contain many of these timing generators as each
can control anywhere from two to all of the electrodes 130
connected to the IPG 102 at a time. However, when there is more
than one timing generator and multiple channels have been actively
programmed then there needs to be a mechanism for suppressing a
second channel from turning on when another is already active.
[0051] The next circuit block contained in the IPG 102 is the
arbitrator. The arbitrator functions by looking at each of the
timing generators' envelope signals and makes sure only one can be
active at a time. If a second tries to activate then the arbitrator
suppresses that signal.
[0052] The arbitrator accomplishes this by bringing each of the
channel envelope signals into a rising edge detection circuit. Once
one is triggered it is fed into the SET pin of an SR flip flop. The
output of this SR-flip flop is fed into all of the other rising
edge detectors in order to suppress them from triggering. The
channel envelope signal is also fed into a falling-edge detector
which is then fed into the RESET of the same SR flip flop. The
output of the SR flip flops are then connected to switches whose
outputs are all tied together that turn on/off that channels
particular biphasic pulse train. Therefore, the output of this
circuit element is a single bi-phasic pulse train and a signal
designating which timing generator that particular pulse train is
sourced from. Essentially, the circuit looks for a channel to go
active. Once it finds one it suppresses all others until that
channel becomes inactive.
[0053] The next section of the circuit works very similarly to the
timing generators to create a high speed burst pulse train that is
then combined with the stimulation pulse train to create a bursted
bi-phasic pulse train if desired.
[0054] It accomplishes this by taking the incoming clock that is
fed from the microcontroller 104 and feeding it into a counter. The
counter can count these rising clock edges infinitely. The counter
is only active during a single phase of the bi-phasic signal and
begins counting as soon as the rising edge is detected. The output
of the counter is fed into a comparator, along with a
microcontroller-programmed register, whose output is connected to
the reset pin on the counter. Therefore, this counter will simply
count to a programmed value and reset. This programmed value is the
burst frequency.
[0055] The output of the comparator is then fed into an edge
detection circuit and then a flip flop that combines it with the
actual stimulation pulse train to create a single phase bursted
stimulation pulse. The entire circuit is duplicated for the second
phase of the signal resulting in the desired bursted bi-phasic
pulse train. The stimulation signal is now handed over to the
electrode logic stage.
[0056] The electrode logic conditions and directs the bi-phasic
signals to the analog section of the ASIC 106. At this point, the
bi-phasic signals contain all of the pertinent timing information,
but none of the required amplitude information. The incoming
signals include the bi-phasic pulse train and another signal
designating which timing generator the current active train came
from. Each electrode logic cell has a register for each timing
generator that stores this particular electrode's 130 amplitude
values for that timing generator. The electrode logic cell uses the
designation signal to determine which register to pull the
amplitude values from, e.g. if the third timing generator is passed
through the arbitration circuit then the electrode logic would read
the value from the third register.
[0057] Once the value is pulled from the register, it goes through
a series of logic gates. The gates first determine that the
electrode 130 should be active. If not, no further action is taken
and the analog section of the electrode output is not activated,
thereby saving precious battery 108 power. Next, a determination is
made if the particular electrode 130 is an anode or cathode. If the
electrode is deemed to be an anode, the electrode logic passes the
amplitude information and the biphasic signal to the positive
current (digital to analog converter) DAC in the analog section of
the ASIC 106. If the electrode is deemed to be a cathode, the
electrode logic passes the amplitude information and the biphasic
signal to the negative current DAC in the analog section of the
ASIC 106. The electrode logic circuit must make these decisions for
each phase of the bi-phasic signal as every electrode 130 will
switch between being an anode and a cathode.
[0058] The analog elements in the ASIC 106 are uniquely designed in
order to produce the desired signals. The basis of analog IC design
is the field effect transistor (FET) and the type of high current
multiple output design required in SCS 100 means that the bulk of
the silicon in the ASIC 106 will be dedicated to the analog
section.
[0059] The signals from the electrode output are fed into each
current DAC when that specific electrode 130 should be activated.
Each electrode 130 has a positive and a negative current DAC,
triggered by the electrode logic and both are never active at the
same time. The job of each current DAC is, when activated, to take
the digital value representing a stimulation current amplitude and
produce an analog representation of this value to be fed into the
output stage. This circuit forms half of the barrier between the
digital and analog sections of the ASIC 106.
[0060] The digital section of the ASIC 106 is built upon a
technology that only allows small voltages to exist. In moving to
the analog section, the output of the current DAC (which is a low
level analog signal) must be amplified to a higher voltage for use
in the analog section. The circuit that performs this task is
called a power level shifter. Because this circuit is built upon
two different manufacturing technologies and requires high
precision analog circuits built upon a digital base, it can be
difficult to implement.
[0061] Once the voltages have been converted for usage in the
analog portion of the ASIC 106 the voltages are passed on to the
output current stages. There are two current sources per electrode
output. One will source a positive current and one will sink a
negative current, but both will never be active simultaneously. The
current sources themselves are made up of analog elements similar
to a Howland current source. There is an input stage, and an
amplification stage with feedback through a sensing component to
maintain the constant current. The input stage takes the analog
voltage values from the power level shifter and produces an output
pulse designated for the amplifier. The amplifier then creates the
pulses of varying voltages but constant current flow. The sources
are capable of sourcing or sinking up to 12.7 mA at 0.1 mA
resolution into a load of up to 1.2 k Ohms. This translates into
range of 15 volts, which will vary depending on the load in order
to keep the current constant.
[0062] The microcontroller 104 to ASIC 106 interface is designed to
be as simple as possible with minimal bus `chatter` in order to
save battery 108 life. The ASIC 106 can be a collection of
registers programmed via a standard I.sup.2C or SPI bus. Since the
ASIC 106 is handling all the power management, there will also be a
power good (PG) line between the two chips 104, 106 in order to let
the microcontroller 104 know when it is safe to power up. The ASIC
106 will also need to use a pin on the microcontroller 104 in order
to generate a hardware interrupt in case anything goes awry in the
ASIC 106. The final connection is the time base for all of the
stimulation circuitry. The ASIC 106 will require two clocks, one
for its internal digital circuitry which will be fed directly from
the microcontroller 104 clock output, and one to base all
stimulation off of which will need to be synthesized by the
microcontroller 104 and fed to the ASIC 106. All commands and
requests to the ASIC 106 will be made over the I.sup.2C or SPI bus
and will involve simply reading a register address or writing to a
register. Even when the ASIC 106 generates a hardware interrupt, it
will be the responsibility of the microcontroller 104 to poll the
ASIC 106 and determine the cause of the interrupt.
[0063] The wireless interface is based upon the FCCs MedRadio
standard operating in the 402-405 MHz range utilizing up to 10
channels for telemetry. The protocol implemented is chosen to
minimize transmission and maximize battery 108 life. All processing
will take place on the user remote/programmer and the only data
transmitted is exactly what will be used in the microcontroller 104
to ASIC 106 bus. That is, all of the wireless packets will contain
necessary overhead information along with only a register address,
data to store in the register, and a command byte instructing the
microcontroller 104 what to do with the data. The overhead section
of the wireless protocol will contain synchronization bits, start
bytes, an address which is synchronized with the IPG's 102 serial
number, and a CRC byte to assure proper transmission. The packet
length is kept as small as possible in order to maintain battery
108 life. Since the IPG 102 cannot listen for packets all the time
due to battery 108 life, it cycles on for a duty cycle of less than
0.05% of the time. This time value can be kept small as long as the
data packets are also small. The user commands needed to run the
system are executed by the entire system using flows.
[0064] The IPG 102 uses an implantable grade Li ion battery 108
with 215 mAHr with zero volt technology. The voltage of the battery
108 at full capacity is 4.1 V and it supplies current only until it
is drained up to 3.3 V which is considered as 100% discharged. The
remaining capacity of the battery 108 can be estimated at any time
by measuring the voltage across the terminals. The maximum charge
rate is 107.5 mA. A Constant Current, Constant Voltage (CCCV) type
of regulation can be applied for faster charging of the battery
108.
[0065] The internal secondary coil 109 is made up of 30 turns of 30
AWG copper magnet wires. The ID, OD, and the thickness of the coil
are 30, 32, and 2 mm, respectively. Inductance L2 is measured to be
58 uH, a 80 nF capacitor is connected to it to make a series
resonance tank at 74 kHz frequency. In the art of induction
charging, two types of rectifiers are considered to convert the
induced AC into usable DC, either a bridge full wave rectifier or a
voltage doubler full wave rectifier. To obtain a higher voltage,
the voltage double full wave rectifier is used in this application.
The rectifier is built with high speed Schottky diodes to improve
its function at high frequencies of the order 100 kHz. A Zener
diode and also a 5V voltage regulator are used for regulation. This
circuit will be able to induce AC voltage, rectify to DC, regulate
to 5V and supply 100 mA current to power management IC that charges
the internal battery 108 by CCCV regulation.
[0066] The regulated 5V 100 mA output from the resonance tank is
fed to, for example, a Power Management Integrated Circuit (PMIC)
MCP73843. This particular chip was specially designed by Microchip
to charge a Li ion battery 108 to 4.1 V by CCCV regulation. The
fast charge current can be regulated by changing a resistor; it is
set to threshold current of 96 mA in the example circuit. The chip
charges the battery 108 to 4.1V as long as the current received is
more than 96 mA. However, if the supply current drops below 96 mA,
it stops to charge the battery 108 until the supply is higher than
96 again. For various practical reasons, if the distance between
the coils increases, the internal secondary coil 109 receives
lesser current than the regulated value, and instead of charging
the battery 108 slowly, it pauses the charging completely until it
receives more than 96 mA. It is understood to those with skill in
the art that other power management chips can be used and the power
management chip is not limited to the PMIC MCP738432 chip.
[0067] All the functions of the IPG 102 are controlled from outside
using a hand held remote controller specially designed for this
device. Along with the remote control, an additional control is
desirable to operate the IPG 102 if the remote control was lost or
damaged. For this purpose a Hall effect based magnet switch was
incorporated to either turn ON or turn OFF the IPG 102 using an
external piece of magnet. Magnet switch acts as a master control
for the IPG 102 to turn on or off. A south pole of sufficient
strength turns the output on and a north pole of sufficient
strength turns the output off. The output is latched so that the
switch continues to hold the state even after the magnet is removed
from its vicinity.
[0068] The IPG 102 is an active medical implant that generates an
electrical signal that stimulates the spinal cord. The signal is
carried through a stimulation lead 140 that plugs directly into the
IPG 102. The IPG 102 recharges wirelessly through an induction coil
109, and communicates via RF radio antenna 110 to change
stimulation parameters. The IPG 102 is implanted up to 3 cm below
the surface of the skin and can be fixed to the fascia by passing
two sutures through holes in the epoxy header 114. The leads 140
are electrically connected to the IPG 102 through a lead contact
system 116, a cylindrical spring-based contact system with
inter-contact silicone seals. The leads 140 are secured to the IPG
102 with a set screw 117 that actuates within locking housing 118.
Set screw compression on the lead's 140 fixation contact can be
governed by a disposable torque wrench. The wireless recharging is
achieved by aligning the exterior induction coil on the charger
with the internal induction coil 109 within the IPG 102. The RF
antenna within the remote's dongle 200 communicates with the RF
antenna 110 in the IPG's 102 epoxy header 114. FIG. 2 illustrates
an exploded view of the IPG 102 assembly.
[0069] The IPG 102 is an assembly of a hermetic titanium (6Al-4V)
casing 120 which houses the battery 108, circuitry 104, 106, and
charging coil 109. The IPG 102 further includes an epoxy header 114
(see FIG. 6), which houses the lead contact assembly 116, locking
housing 118, and RF antenna 110 (see FIGS. 6 and 7). The internal
electronics are connected to the components within the epoxy head
through a hermetic feedthrough 122, as shown in FIG. 3. The
feedthrough 122 is a titanium (6Al-4V) flange with an alumina
window and gold trimming. Within the alumina window are thirty-four
platinum-iridium (90-10) pins that interface internally with a
direct solder to the circuit board, and externally with a series of
platinum iridium wires laser-welded to the antenna 110 and lead
contacts 126. The IPG 102 interfaces with 32 electrical contacts
126, which are arranged in four rows of eight contacts 126. Thirty
two of the feedthrough's 122 pins 124 interface with the contacts
126, while two interface with the antenna 110, one to the ground
plane and one to the antenna 110 feed.
[0070] FIGS. 4 and 5 depict a lead contact system 115 and assembly
116, respectively. The lead contacts 126 consist of an MP35N
housing 128 with a platinum-iridium 90-10 spring 129. Each contact
126 is separated by a silicone seal 127. At the proximal end of
each stack of 8 contacts 126 is a titanium (6Al-4V) cap 125 which
acts as a stop for the lead 140. At the distal end is a titanium
(6Al-4V) set screw 119 and block 118 for lead fixation. At the lead
entrance point is a silicone tube 123 which provides strain relief
as the lead 140 exits the head unit 114, and above the set screw
119 another silicone tube 131 with a small internal canal allows
the torque wrench to enter but does not allow the set screw 119 to
back out. In addition to the contacts 126 and antenna 110, the
header 114 also contains a radiopaque titanium (6Al-4V) tag 132
which allows for identification of the device under fluoroscopy.
The overmold of the header 114 is Epotek 301, a two-part,
biocompatible epoxy. FIGS. 4, 5, 6, and 7 depict illustrations of
lead contact system 115, lead contact assembly 116, head unit
assembly 114, and RF antenna 110, respectively.
[0071] Internal to the titanium (6Al-4V) case 120 are the circuit
board 105, battery 108, charging coil 109, and internal plastic
support frame. The circuit board 105 can be a multi-layered FR-4
board with copper traces and solder mask coating. Non-solder masked
areas of the board can be electroless nickel immersion gold. The
implantable battery 108, all surface mount components, ASIC 106,
microcontroller 104, charging coil 109, and feedthrough 122 will be
soldered to the circuit board 105. The plastic frame, made of
either polycarbonate or ABS, will maintain the battery's 108
position and provide a snug fit between the circuitry 105 and case
120 to prevent movement. The charging coil 109 is a wound coated
copper.
Leads
[0072] The percutaneous stimulation leads 140, as depicted in FIG.
8, are a fully implantable electrical medical accessory to be used
in conjunction with the implantable SCS 100. The primary function
of the lead is to carry electrical signals from the IPG 102 to the
target stimulation area on the spinal cord. Percutaneous
stimulation leads 140 provide circumferential stimulation. The
percutaneous stimulation leads 140 provide a robust, flexible, and
bio-compatible electric connection between the IPG 102 and
stimulation area. The leads 140 are surgically implanted through a
spinal needle, or epidural needle, and are driven through the
spinal canal using a steering stylet that passes through the center
of the lead 140. The leads 140 are secured mechanically to the
patient using either an anchor or a suture passed through tissue
and tied around the body of the lead 140. The leads 140 are secured
at the proximal end with a set-screw 119 on the IPG 102 which
applies radial pressure to a blank contact on the distal end of the
proximal contacts.
[0073] The percutaneous stimulation leads 140 consist of a
combination of implantable materials. Stimulation electrodes 130 at
the distal end and electrical contacts at the proximal end are made
of a 90-10 platinum-iridium alloy. This alloy is utilized for its
bio-compatibility and electrical conductivity. The electrodes 130
are geometrically cylindrical. The polymeric body of the lead 140
is polyurethane, chosen for its bio-compatibility, flexibility, and
high lubricity to decrease friction while being passed through
tissue. The polyurethane tubing has a multi-lumen cross section,
with one center lumen 142 and eight outer lumens 144. The center
lumen 142 acts as a canal to contain the steering stylet during
implantation, while the outer lumens 144 provide electrical and
mechanical separation between the wires 146 that carry stimulation
from the proximal contacts to distal electrodes 130. These wires
146 are a bundle of MP35N strands with a 28% silver core. The wires
146 are individually coated with ethylene tetrafluoroethylene
(ETFE), to provide an additional non-conductive barrier. The wires
146 are laser welded to the contacts and electrodes 130, creating
an electrical connection between respective contacts on the
proximal and distal ends. The leads 140 employ a platinum-iridium
plug 148, molded into the distal tip of the center lumen 142 to
prevent the tip of the steering stylet from puncturing the distal
tip of the lead 140. Leads 140 are available in a variety of 4 and
8 electrode 130 configurations. These leads 140 have 4 and 8
proximal contacts (+1 fixation contact), respectively.
Configurations vary by electrode 130 number, electrode 130 spacing,
electrode 130 length, and overall lead 140 length.
[0074] The paddle stimulation leads 141, as depicted in FIG. 9, are
a fully implantable electrical medical accessory to be used in
conjunction with the implantable SCS 100. The primary function of
the paddle lead 141 is to carry electrical signals from the IPG 102
to the target stimulation area on the spinal cord. The paddle leads
141 provide uni-direction stimulation across a 2-dimensional array
of electrodes 130, allowing for greater precision in targeting
stimulation zones. The paddle stimulation leads 141 provide a
robust, flexible, and bio-compatible electric connection between
the IPG 102 and stimulation area. The leads 141 are surgically
implanted through a small incision, usually in conjunction with a
laminotomy or laminectomy, and are positioned using forceps or a
similar surgical tool. The leads 141 are secured mechanically to
the patient using either an anchor or a suture passed through
tissue and tied around the body of the lead 141. The leads 141 are
secured at the proximal end with a set-screw on the IPG 102 which
applies radial pressure to a fixation contact on the distal end of
the proximal contacts.
[0075] The paddle stimulation leads 141 consist of a combination of
implantable materials. Stimulation electrodes 130 at the distal end
and electrical contacts at the proximal end are made of a 90-10
platinum-iridium alloy utilized for its bio-compatibility and
electrical conductivity. The polymeric body of the lead 141 is
polyurethane, chosen for its bio-compatibility, flexibility, and
high lubricity to decrease friction while being passed through
tissue. The polyurethane tubing has a multi-lumen cross section,
with one center lumen 142 and eight outer lumens 144. The center
lumen 142 acts as a canal to contain the steering stylet during
implantation, while the outer lumens 144 provide electrical and
mechanical separation between the wires 146 that carry stimulation
from the proximal contacts to distal electrodes 130. These wires
146 are a bundle of MP35N strands with a 28% silver core. The wires
146 are individually coated with ethylene tetrafluoroethylene
(ETFE), to provide an additional non-conductive barrier. At the
distal tip of the paddle leads 141 is a 2-dimensional array of flat
rectangular electrodes 130 molded into a flat silicone body 149. In
an embodiment, one side of the rectangular electrodes 130 is
exposed, providing uni-directional stimulation. The wires 146 are
laser welded to the contacts and electrodes 130, creating an
electrical connection between respective contacts on the proximal
and distal ends. Also molded into the distal silicone paddle is a
polyester mesh 147 adding stability to the molded body 149 while
improving aesthetics by covering wire 146 routing. The number of
individual 8-contact leads 141 used for each paddle 141 is governed
by the number of electrodes 130. Electrodes 130 per paddle 141
range from 8 to 32, split into between one and four proximal lead
141 ends. Each proximal lead 141 has 8 contacts (+1 fixation
contact). Configurations vary by electrode 130 number, electrode
130 spacing, electrode length, and overall lead length.
[0076] The lead extensions 150, as depicted in FIG. 10, are a fully
implantable electrical medical accessory to be used in conjunction
with the implantable SCS 100 and either percutaneous 140 or paddle
141 leads. The primary function of the lead extension 150 is to
increase the overall length of the lead 140, 141 by carrying
electrical signals from the IPG 102 to the proximal end of the
stimulation lead 140, 141. This extends the overall range of the
lead 140, 141 in cases where the length of the provided leads 140,
141 are insufficient. The lead extensions 150 provide a robust,
flexible, and bio-compatible electric connection between the IPG
102 and proximal end of the stimulation lead 140, 141. The
extensions 150 may be secured mechanically to the patient using
either an anchor or a suture passed through tissue and tied around
the body of the extension 150. Extensions 150 are secured at the
proximal end with a set-screw 119 on the IPG 102 which applies
radial pressure to a fixation contact on the distal end of the
proximal contacts of the extension 150. The stimulation lead 140,
141 is secured to the extension 150 in a similar fashion, using a
set screw 152 inside the molded tip of extension 150 to apply a
radial pressure to the fixation contact at the proximal end of the
stimulation lead 140, 141.
[0077] The lead extension 150 consists of a combination of
implantable materials. At the distal tip of the extension 150 is a
1.times.8 array of implantable electrical contacts 154, each
consisting of MP35 housing 128 and 90-10 platinum-iridium spring. A
silicone seal 127 separates each of the housings 128. At the
proximal end of the contacts is a titanium (6Al4V) cap which acts
as a stop for the lead, and at the distal tip, a titanium (6Al4V)
block and set screw 152 for lead fixation. The electrical contacts
at the proximal end are made of a 90-10 platinum-iridium alloy
utilized for its bio-compatibility and electrical conductivity. The
polymeric body 156 of the lead 150 is polyurethane, chosen for its
bio-compatibility, flexibility, and high lubricity to decrease
friction while being passed through tissue. The polyurethane tubing
158 has a multi-lumen cross section, with one center lumen 142 and
eight outer lumens 144. The center lumen 142 acts as a canal to
contain the steering stylet during implantation, while the outer
lumens 144 provide electrical and mechanical separation between the
wires 146 that carry stimulation from the proximal contacts to
distal electrodes. These wires 146 are a bundle of MP35N strands
with a 28% silver core. The wires 146 are individually coated with
ethylene tetrafluoroethylene (ETFE), to provide an additional
non-conductive barrier. Each lead extension 150 has 8 proximal
cylindrical contacts (+1 fixation contact).
[0078] The lead splitter 160, as depicted in FIG. 11, is a fully
implantable electrical medical accessory which is used in
conjunction with the SCS 100 and typically a pair of 4-contact
percutaneous leads 140. The primary function of the lead splitter
160 is to split a single lead 140 of eight contacts into a pair of
4 contact leads 140. The splitter 160 carries electrical signals
from the IPG 102 to the proximal end of two 4-contact percutaneous
stimulation leads 140. This allows the surgeon access to more
stimulation areas by increasing the number of stimulation leads 140
available. The lead splitter 160 provides a robust, flexible, and
bio-compatible electrical connection between the IPG 102 and
proximal ends of the stimulation leads 140. The splitters 160 may
be secured mechanically to the patient using either an anchor or a
suture passed through tissue and tied around the body of the
splitter 160. Splitters 160 are secured at the proximal end with a
set-screw 119 on the IPG 102 which applies radial pressure to a
fixation contact on the distal end of the proximal contacts of the
splitter 160. The stimulation leads 140 are secured to the splitter
160 in a similar fashion, using a pair of set screws inside the
molded tip of splitter 160 to apply a radial pressure to the
fixation contact at the proximal end of each stimulation lead
140.
[0079] The lead splitter 160 consists of a combination of
implantable materials. At the distal tip of the splitter 160 is a
2.times.4 array of implantable electrical contacts 162, with each
contact 162 consisting of MP35 housing 128 and 90-10
platinum-iridium spring. A silicone seal 127 separates each of the
housings 128. At the proximal end of each row of contacts 162 is a
titanium (6Al4V) cap which acts as a stop for the lead, and at the
distal tip, a titanium (6Al4V) block and set screw for lead
fixation. The electrical contacts at the proximal end of the
splitter 160 are made of a 90-10 platinum-iridium alloy utilized
for its bio-compatibility and electrical conductivity. The
polymeric body 164 of the lead 160 is polyurethane, chosen for its
bio-compatibility, flexibility, and high lubricity to decrease
friction while being passed through tissue. The polyurethane tubing
166 has a multi-lumen cross section, with one center lumen 142 and
eight outer lumens 144. The center lumen 142 acts as a canal to
contain the steering stylet during implantation, while the outer
lumens 144 provide electrical and mechanical separation between the
wires 146 that carry stimulation from the proximal contacts to
distal electrodes 130. These wires 146 are a bundle of MP35N
strands with a 28% silver core. The wires 146 are individually
coated with ethylene tetrafluoroethylene (ETFE), to provide an
additional non-conductive barrier. Each lead splitter 160 has 8
proximal contacts (+1 fixation contact), and 2 rows of 4 contacts
162 at the distal end.
Anchors
[0080] The lead anchor 170, as depicted in FIGS. 12 and 13, is a
fully implantable electrical medical accessory which is used in
conjunction with both percutaneous 140 and paddle 141 stimulation
leads. The primary function of the lead anchor 170 is to prevent
migration of the distal tip of the lead 140, 141 by mechanically
locking the lead 140, 141 to the tissue. There are currently two
types of anchors 170, a simple sleeve 171, depicted in FIG. 12, and
a locking mechanism 172, depicted in FIG. 13, and each has a
slightly different interface. For the simple sleeve type anchor
171, the lead 140, 141 is passed through the center thru-hole 174
of the anchor 171, and then a suture is passed around the outside
of the anchor 171 and tightened to secure the lead 140, 141 within
the anchor 171. The anchor 171 can then be sutured to the fascia.
The locking anchor 172 uses a set screw 176 for locking purposes,
and a bi-directional disposable torque wrench for locking and
unlocking. Tactile and audible feedback is provided for both
locking and unlocking.
[0081] Both anchors 171, 172 can be molded from implant-grade
silicone, but the locking anchor 172 uses an internal titanium
assembly for locking. The 3-part mechanism is made of a housing
175, a locking set screw 176, and a blocking set screw 177 to
prevent the locking set screw from back out. All three components
can be titanium (6Al4V). The bi-directional torque wrench can have
a plastic body and stainless steel hex shaft.
Wireless Dongle
[0082] The wireless dongle 200 is the hardware connection to a
smartphone/mobile 202 or tablet 204 that allows communication
between the trial generator 107 or IPG 102 and the
smartphone/mobile device 202 or tablet 204, as illustrated in FIG.
14. During the trial or permanent implant phases, the wireless
dongle 200 is connected to the tablet 204 through the tablet 204
specific connection pins and the clinician programmer software on
the tablet 204 is used to control the stimulation parameters. The
commands from the clinician programmer software are transferred to
the wireless dongle 200 which is then transferred from the wireless
dongle 200 using RF signals to the trial generator 107 or the IPG
102. Once the parameters on the clinician programmers have been
set, the parameters are saved on the tablet 204 and can be
transferred to the patient programmer software on the
smartphone/mobile device 202. The wireless dongle 200 is composed
of an antenna, a microcontroller (having the same specifications as
the IPG 102 and trial generator 107), and a pin connector to
connect with the smartphone/mobile device 202 and the tablet
204.
Charger
[0083] The IPG 102 has a rechargeable lithium ion battery 108 to
power its activities. An external induction type charger 210 (FIG.
1) wirelessly recharges the included battery 108 inside the IPG
102. The charger 210 is packaged into a housing and consists of a
rechargeable battery, a primary coil of wire and a printed circuit
board (PCB) containing the electronics. In operation, charger 210
produces a magnetic field and induces voltage into the secondary
coil 109 in the IPG 102. The induced voltage is then rectified and
used to charge the battery 108 inside the IPG 102. To maximize the
coupling between the coils, both internal and external coils are
combined with capacitors to make them resonate at a particular
common frequency. The coil acting as an inductor L forms an LC
resonance tank. The charger uses a Class-E amplifier topology to
produce the alternating current in the primary coil around the
resonant frequency. The charger 210 features include, but are not
limited to: [0084] Charge IPG 102 wirelessly [0085] Charge up to a
maximum depth of 30 mm [0086] Integrated alignment sensor indicates
alignment between the charger and IPG 102 resulting in higher power
transfer efficiency [0087] Alignment sensor provides audible and
visual feedback to the user [0088] Compact and Portable
[0089] A protected type of cylindrical Li ion battery is used as
the charger 210 battery. A Class-E power amplifier topology is a
much used type of amplifier for induction chargers, especially for
implantable electronic medical devices. Due to the Class-E power
amplifier's relatively high theoretical efficiency it is often used
for devices where high efficiency power transfer is necessary. A
0.1 ohm high wattage resistor is used in series to sense the
current through this circuit.
[0090] The primary coil L1 is made by 60 turns of Litz wire type
100/44-100 strands of 44 AWG each. The Litz wire solves the problem
of skin effect and keeps its impedance low at high frequencies.
Inductance of this coil was initially set at 181 uH, but backing it
with a Ferrite plate increases the inductance to 229.7 uH. The
attached ferrite plate focuses the produced magnetic field towards
the direction of the implant. Such a setup helps the secondary coil
receive more magnetic fields and aids it to induce higher
power.
When the switch is ON, the resonance is at frequency
f = 1 2 .pi. L 1 C 2 ##EQU00001##
When the switch is OFF, it shifts to
f = 1 2 .pi. L 1 C 1 C 2 C 1 + C 2 ##EQU00002##
In a continuous operation the resonance frequency will be in the
range
1 2 .pi. L 1 C 2 < f < 1 2 .pi. L 1 C 1 C 2 C 1 + C 2
##EQU00003##
[0091] To make the ON and OFF resonance frequencies closer, a
relatively larger value of C1 can be chosen by a simple criteria as
follows
C1=nC2; a value of n=4 was used in the example above; in most cases
3<n<10.
[0092] The voltages in these Class-E amplifiers typically go up to
the order of 300 VAC. Capacitors selected must be able to withstand
these high voltages, sustain high currents and still maintain low
Effective Series Resistance (ESR). Higher ESRs result in
unnecessary power losses in the form of heat. The circuit is
connected to the battery through an inductor which acts as a choke.
The choke helps to smoothen the supply to the circuit. The N
Channel MOSFET acts as a switch in this Class-E power amplifier. A
FET with low ON resistance and with high drain current I.sub.d is
desirable.
[0093] In summary, the circuit is able to recharge the IPG 102
battery 108 from 0 to 100% in approximately two hours forty-five
minutes with distance between the coils being 29 mm. The primary
coil and the Class-E amplifier draws DC current of 0.866 A to
achieve this task. To improve the efficiency of the circuit, a
feedback closed loop control is implemented to reduce the losses.
The losses are minimum when the MOSFET is switched ON and when the
voltage on its drain side is close to zero.
[0094] The controller takes the outputs from operational
amplifiers, checks if the outputs meet the criteria, then triggers
the driver to switch ON the MOSFET for the next cycle. The
controller can use a delay timer, an OR gate and a 555 timer in
monostable configuration to condition the signal for driver. When
the device is switched ON, the circuit will not function right away
as there is no active feedback loop. The feedback becomes active
when the circuit starts to function. To provide an active feedback
loop, an initial external trigger is applied to jump start the
system.
Alignment Sensor
[0095] The efficiency of the power transfer between the external
charger 210 and the internal IPG 102 will be maximum only when the
charger 210 and IPG 102 are properly aligned. An alignment sensor
is provided to ensure proper alignment as part of the external
circuit design, and is based on the principle of reflected
impedance. When the external coil is brought closer to the internal
coil, the impedance of both circuits change. The sensing is based
on measuring the reflected impedance and testing whether it crosses
the threshold. A beeper provides an audible feedback to the patient
and a LED provides visual feedback.
[0096] When the impedance of the circuit changes, the current
passing through it also changes. A high power 0.1 ohm resistor can
be used in the series of the circuit to monitor the change in
current. The voltage drop across the resistor is amplified 40 times
and then compared to a fixed threshold value using an operational
amplifier voltage comparator. The output is fed to a timer chip
which in turn activates the beeper and LED to provide feedback to
the user.
[0097] The circuit can sense the alignment up to a distance of
approximately 30 mm. The current fluctuation in the circuit depends
on more factors than reflected impedance alone and the circuit is
sensitive to other parameters of the circuit as well. To reduce the
sensitivity related to other parameters, one option is to eliminate
interference of all the other factors and improve the functionality
of the reflected impedance sensor--which is very challenging to
implement within the limited space available for circuitry. Another
option is to use a dedicated sensor chip to measure the reflected
impedance.
[0098] A second design uses sensors designed for proximity detector
or metal detectors for alignment sensing. Chips designed to detect
metal bodies by the effect of Eddy currents on the HF losses of a
coil can be used for this application. The TDE0160 is an example of
such a chip.
[0099] The external charger is designed to work at 75 to 80 kHz,
whereas the proximity sensor was designed for 1 MHz. The sensor
circuit is designed to be compatible with the rest of the external
and is fine tuned to detect the internal IPG 102 from a distance of
approximately 30 mm.
Programmer
[0100] The Clinician Programmer is an application that is installed
on a tablet 204. It is used by the clinician to set the stimulation
parameters on the trial generator 107 or IPG 102 during trial and
permanent implantation in the operating room. The clinician
programmer is capable of saving multiple settings for multiple
patients and can be used to adjust the stimulation parameters
outside of the operations room. It is capable of changing the
stimulation parameters though the RF wireless dongle 200 when the
trial generator 107 or IPG 102 which has been implanted in the
patient is within the RF range. In addition, it is also capable of
setting or changing the stimulation parameters on the trial
generator 107 and/or the IPG 102 through the internet when both the
tablet 204 and the Patient Programmers on a smartphone/mobile
device 202 both have access to the internet.
[0101] The Patient Programmer is an application that is installed
on a smartphone/mobile device 202. It is used by the patient to set
the stimulation parameters on the trial generator 107 or IPG 102
after trial and permanent implantation outside the operating room.
The clinician programmer is capable of saving multiple settings for
multiple patients and can be transferred to the Patient Programmer
wirelessly when the Clinician Programmer tablet 204 and the Patient
Programmer smartphone/mobile device 202 are within wireless range
such as Bluetooth from each other. In the scenario where the
Clinician Programmer tablet 204 and the Patient Programmer
smartphone/mobile device 202 are out of wireless range from each
other, the data can be transferred through the internet where both
devices 202, 204 have wireless access such as Wi-Fi. The Patient
Programmer is capable of changing the stimulation parameters on the
trial generator 107 or IPG 102 though the RF wireless dongle 200
when the trial generator 107 or IPG implanted in the patient is
within the RF range.
Tuohy Needle
[0102] The Tuohy needle 240, as depicted in FIG. 15, is used in
conjunction with a saline-loaded syringe for loss-of-resistance
needle placement, and percutaneous stimulation leads 140, for lead
140 placement into the spinal canal. The Tuohy epidural needle 240
is inserted slowly into the spinal canal using a loss-of-resistance
technique to gauge needle 240 depth. Once inserted to the
appropriate depth, the percutaneous stimulation lead 140 is passed
through the needle 240 and into the spinal canal.
[0103] The epidural needle 240 is a non-coring 14G stainless steel
spinal needle 240 and will be available in lengths of 5'' (127 mm)
and 6'' (152.4). The distal tip 242 of the needle 240 has a slight
curve to direct the stimulation lead 140 into the spinal canal. The
proximal end 246 is a standard Leur-Lock connection 248.
Stylet
[0104] The stylet 250, as depicted in FIG. 16, is used to drive the
tip of a percutaneous stimulation lead 140 to the desired
stimulation zone by adding rigidity and steerability. The stylet
250 wire 252 passes through the center lumen 142 of the
percutaneous lead 140 and stops at the blocking plug at the distal
tip of the lead 140. The tip of the stylet 250 comes with both
straight and curved tips. A small handle 254 is used at the
proximal end of the stylet 250 to rotate the stylet 250 within the
center lumen 142 to assist with driving. This handle 254 can be
removed and reattached allowing anchors 170 to pass over the lead
140 while the stylet 250 is still in place. The stylet 250 wire 252
is a PTFE coated stainless steel wire and the handle 254 is
plastic.
Passing Elevator
[0105] The passing elevator 260, as depicted in FIG. 17, is used
prior to paddle lead 141 placement to clear out tissue in the
spinal canal and help the surgeon size the lead to the anatomy. The
passing elevator 260 provides a flexible paddle-shaped tip 262 to
clear the spinal canal of obstructions. The flexible tip is
attached to a surgical handle 264.
[0106] The passing elevator 260 is a one-piece disposable plastic
instrument made of a flexible high strength material with high
lubricity. The flexibility allows the instrument to easily conform
to the angle of the spinal canal and the lubricity allows the
instrument to easily pass through tissue.
Tunneling Tool
[0107] The tunneling tool 270, as depicted in FIG. 18, is used to
provide a subcutaneous canal to pass stimulation leads 140 from the
entrance point into the spinal canal to the IPG implantation site.
The tunneling tool 270 is a long skewer-shaped tool with a ringlet
handle 272 at the proximal end 274. The tool 270 is covered by a
plastic sheath 276 with a tapered tip 278 which allows the tool 270
to easily pass through tissue. Once the IPG 102 implantation zone
is bridge to the lead 140 entrance point into the spinal canal, the
inner core 275 is removed, leaving the sheath 276 behind. The leads
140 can then be passed through the sheath 276 to the IPG 102
implantation site. The tunneling tool 270 is often bent to assist
in steering through the tissue.
[0108] The tunneling tool 270 is made of a 304 stainless steel core
with a fluorinated ethylene propylene (FEP) sheath 276. The 304
stainless steel is used for its strength and ductility during
bending, and the sheath 276 is used for its strength and
lubricity.
Torque Wrench
[0109] The torque wrench 280, as depicted in FIG. 19, is used in
conjunction with the IPG 102, lead extension 150 and lead splitter
160 to tighten the internal set screw 119, which provides a radial
force against the fixation contact of the stimulation leads 140,
141, preventing the leads 140, 141 from detaching. The torque
wrench 280 is also used to lock and unlock the anchor 170. The
torque wrench 280 is a small, disposable, medical instrument that
is used in every SCS 100 case. The torque wrench 280 provides
audible and tactile feedback to the surgeon that the lead 140, 141
is secured to the IPG 102, extension 150, or splitter 160, or that
the anchor 170 is in the locked or unlocked position.
[0110] The torque wrench 280 is a 0.9 mm stainless steel hex shaft
282 assembled with a plastic body 284. The wrench's 280 torque
rating is bi-directional, primarily to provide feedback that the
anchor 170 is either locked or unlocked. The torque rating allows
firm fixation of the set screws 119, 152 against the stimulation
leads 140, 141 without over-tightening.
Trial Patch
[0111] The trial patch is used in conjunction with the trialing
pulse generator 107 to provide a clean, ergonomic protective cover
of the stimulation lead 140, 141 entrance point in the spinal
canal. The patch is also intended to cover and contain the trial
generator 107. The patch is a large, adhesive bandage that is
applied to the patient post-operatively during the trialing stage.
The patch completely covers the leads 140, 141 and generator 107,
and fixates to the patient with anti-microbial adhesive.
[0112] The patch is a watertight, 150 mm.times.250 mm
anti-microbial adhesive patch. The watertight patch allows patients
to shower during the trialing period, and the anti-microbial
adhesive decreases the risk of infection. The patch will be made of
polyethylene, silicone, urethane, acrylate, and rayon.
Magnetic Switch
[0113] The magnetic switch is a magnet the size of a coin that,
when placed near the IPG 102, can switch it on or off. The
direction the magnet is facing the IPG 102 determines if the
magnetic switch is switching the IPG 102 on or off.
[0114] As shown in FIG. 20, the implantable pulse generator (IPG)
102 includes an RF transceiver module 103, a processor such as the
microcontroller 104 and a programmable signal generator such as the
ASIC 106. The transceiver module 103 manages wireless communication
between the microcontroller 104 and external remote (e.g., dongle
200 connected to either the smartphone/mobile 202 or tablet
204).
[0115] In the embodiment shown in FIG. 20, a treatment control
module 14 stored in a flash memory 13 of the microcontroller 104 is
executed by the microcontroller to centrally control operation of
every component and circuits of the IPG 102 with the exception of
an independently operated charger (not shown) that charges the
battery 108. Specifically, the treatment control module 14 handles
programming of the RF transceiver module 103 and signal generator
106 among other functions. Communications among the microcontroller
104, RF transceiver module 103 and signal generator 106 are
performed over a bus 16 such as the Serial Peripheral Interface
(SPI) bus.
[0116] One exemplary microcontroller 104 may be MSP430F5328 from
Texas Instruments of Dallas, Tex. as it has very low power usage,
large amount of memory, integrated peripherals and small physical
size.
[0117] As shown in more detail in FIG. 21, the signal generator 106
includes memory (control registers 18), timing generator 20,
arbitrator circuit 22, high frequency generator 24, electrode
driver 26 which are all coupled to each other. All components in
FIG. 21 have access to and are supplied with signal parameters
stored in the control registers 18 through a register bus 28.
[0118] One of the many novel features of the IPG 102 is that the
control registers 18 in the signal generator 106 have sufficient
memory to store all of the signal parameters necessary to drive the
electrodes E1-E32 independently of the microcontroller 104. As a
result, the microcontroller 104 can be placed in a standby mode
once it programs all of the pulse parameters in the control
registers 18 and the treatment control module 14 instructs the
signal generator to generate the stimulation signals by setting the
stimulation enable pin STIM-EN to logic high. In the embodiment
shown in which the microcontroller is MSP430F5328, the treatment
control module 14 places the microcontroller in LPM3 Standby Mode.
In an LPM3 mode, the master clock (main clock) that drives the
instruction execution unit of the microcontroller 104 is turned
off, essentially turning the microcontroller off so as to conserve
battery power.
[0119] The microcontroller can be waken up from the standby mode by
an interrupt signal IRQ which can be transmitted by the transceiver
module 103 when it receives an appropriate instruction from a
remote control device.
[0120] As discussed earlier, in the IPG 102, the electrodes E1-E32
may be grouped into stimulation sets (stimulation channels). Each
stimulation channel represents one particular stimulation
signal/pattern which is applied to the associated electrodes. In
the embodiment shown, the IPG 102 can accommodate up to 16 channels
(ch1 through ch16). Each electrode can belong to one or more
channels up to the maximum number of channels and each channel can
be associated with at least 2 electrodes to a maximum of 32
electrodes. Accordingly, one electrode can belong to all 16
channels.
[0121] For example, as shown in FIG. 25, channel 1 includes
electrodes E1 and E2, channel 2 includes electrodes E2 and E3 while
channel 3 includes electrodes E2-E4 and E6-E8. Thus, electrode E2
belongs to channels 1, 2 and 3, electrode E3 belongs to channels 2
and 3, while electrodes E4 and E5-E8 belong to only channel 3 and
electrode E1 belong to only channel 1. In FIG. 25, electrode E5 is
unused and therefore does not belong to any channel. Data that
associates electrodes to particular stimulation channels are stored
in the control registers 18.
[0122] As will be discussed more fully later herein, for each
channel, the IPG 102 is capable of programming the amplitude,
frequency and duration of both the first phase pulse (pulse1) and
the second phase pulse (pulse2) of a biphasic pulse (see FIG. 26,
for example) for all of the stimulation channels. As seen in FIG.
26, waveforms of a biphasic pulse comprising pulse1 and pulse2 of
an active channel (ch_act) are combined into a single pulse pattern
which are applied to the electrodes belonging to the active channel
such as E1 and E2. FIG. 26 illustrates the pulse width and
amplitude variations. As can be seen, the width of pulse1 is wider
than that of pulse2 while the amplitude for pulse1 as scaled by the
pulse scaler 38 (see positive pulse of E1, for example) is lower
than that of pulse2 (see negative pulse of E1, for example).
[0123] The pulse parameters for pulse1 are independent of those for
pulse2 for maximum flexibility in managing pain. As an example and
referring to FIG. 25, assume that the current path between
electrodes E1 and E2 (channel 1) affects a nerve path to a left leg
while the current path between electrodes E2 and E3 (channel 2)
affects a nerve path to a right leg. When the patient complains of
more pain on the right leg than the left leg, a physician
programming the IPG 102 may associate electrodes E1 and E2 to
channel 1 with a stimulation pulse pattern having 100 Hz in
frequency and associate electrodes E2 and E3 to channel 2 with a
pulse pattern having 1000 Hz in frequency and a higher current
amplitude than channel 1. In this way, smaller current is applied
to channel 1 for the left leg and higher current is applied to
channel 2 for the right leg. Advantageously, the flexibility of the
IPG 102 allows just the right amount of current to each affected
area of the patient.
[0124] The control registers 18 include standard read/write
registers 18 that can be accessed by the SPI bus 16 and register
bus 28. The control registers 18 are configured as an array of
8-bit registers, each with a unique address. The control registers
18 are programmed by the microcontroller 104 to store all pulse
parameters that are necessary for the signal generator 106 to
generate all of the stimulation channel patterns without any
intervention from the microcontroller. The pulse parameters include
stimulation channel timing settings, current (pulse amplitude)
scaler settings, calibration data and electrode group
parameters.
[0125] For each channel, the control registers 18 store the rising
and falling edges of the channel itself, rising and falling edges
of each of the two pulses (pulse1 and pulse2), period of the
biphasic pulse, active channel period (channel envelope), and
current scaling (pulse amplitude) values for both pulses (pulse1
and pulse2). For each channel, the control registers 18 also store
burst frequency data (as will be explained later herein) such as
burst period for both pulses (pulse1 and pulse2). For each channel,
the control registers 18 also store data regarding which electrodes
E1-E32 belong to that channel. For each channel and for each
electrode within that channel, the control registers 18 store
source/sink data for both pulses of the biphasic pulse (pulse1 and
pulse2), i.e., whether each electrode will be sourcing current or
sinking current during pulse1 and pulse2. All parameters are
specified with reference to the origin and is in units of
microseconds.
[0126] The timing generator 20 generates stimulation timing signals
which comprise pulse1, pulse2, and a channel pulse "ch" (channel
envelope waveform) for all 16 channels based on the pulse timing
parameters stored in the control registers 18. If all 16 channels
are programmed by the clinician, then the timing generator 20
generates the pulse1, pulse2, and channel pulse data ch for all 16
channels simultaneously.
[0127] As an example, pulse1 and pulse2 waveforms of FIG. 26
illustrate the output waveforms from the timing generator 20 for an
exemplary biphasic pulse of a particular channel. The channel
envelope data ch_act defines the start and end of an active portion
of each channel as well as the channel period, which can be defined
as the time between two adjacent rising edges of the channel.
Pulse1 and Pulse2 define the start and end of each of the two
phases of the biphasic pulse.
[0128] Because of the flexibility of the IPG 102, more than one
channel could become active at any time when stimulation patterns
of multiple channels are programmed. The arbitrator 22 is designed
to resolve the overlapping channel (channel contention) problem by
ensuring that only one channel is active at any one time. Among
others, the circuits in the arbitrator 22 are designed with two
rules. The first rule is that when an active channel is being
selected (i.e., a channel currently in progress), all other
channels attempting to go active are suppressed and discarded. The
second rule is that when two or more channels are about to become
active with simultaneous rising edges in ch, an active channel will
be determined based on a predetermined channel priority.
[0129] In the embodiment shown, the arbitrator 22 has been
programmed such that the lowest numbered channel will be given
priority and the remaining simultaneous channels will be discarded.
Since there are 16 channels (ch1 through ch16) in the IPG 102,
channel one has the highest priority while channel 16 has the
lowest priority.
[0130] The output of the arbitrator 22 includes pulse timing signal
p1_act and p2_act which are the same waveforms as pulse1 and pulse2
of an active channel. The arbitrator 22 also outputs the channel
envelope of an active channel (ch_act as shown in FIG. 26, for
example) for use by the high frequency generator 24 as well as the
channel number of the active channel (ch_code), which will be used
by the electrode driver 26, as will be explained later herein. In
the embodiment shown, the channel number is a 4-bit code that
identifies the number of the active channel. For example, `0001`
represents channel 2 while `1111` represents channel 16.
[0131] FIG. 27 provides an example of the channel arbitration by
the arbitrator 22. In FIG. 27, channel number one has 3 electrodes
E1-E3 and channel number two has 2 electrodes E1-E2. As the channel
period for the two channels is different, they will overlap from
time to time. In the illustration, channel one is active
(ch1_active) when channel two attempts to become active. At that
time, the arbitrator executes the first rule, and will suppress
channel two and prevent it from becoming active. Thus, only the
pulse1/pulse2 signals that drive the electrodes for channel one
(active channel) will be output by the arbitrator 22. The
pulse1/pulse2 signals and channel envelope signal for channel two
(ch2_active) are shown in dotted lines to show that they have been
suppressed by the arbitrator 22.
[0132] The high frequency generator 24 receives the p1_act and
p2_act waveforms from the arbitrator 22, decides whether to
modulate the received signals based on the stored parameters in the
control registers 18. If the decision is no, then the high
frequency generator 24 passes the received pulse signals unaltered
to the electrode driver 26.
[0133] If the decision is a yes, however, the high frequency
generator 24 modulates the received signals at a burst frequency
that has been programmed into the control registers 18. The burst
frequency is higher than the frequency of the received signals
p1_act and p2_act.
[0134] The electrode driver 26 receives the output (p1, p2 and
ch_code) of the high frequency generator 24, amplifies the received
signal according to the pulse amplitude parameters stored in the
control registers 18, and outputs the final stimulation pattern for
each channel to be applied through the electrodes E1-E32. As
discussed above, the burst pulse parameters stored in the control
registers 18 have separate frequency values for pulse1 and pulse2
such that an asymmetric pulse shape with positive and negative
pulses having different frequency values can be generated by the
electrode driver 26.
[0135] FIG. 22 is a more detailed functional block diagram of the
high frequency generator of FIG. 21. The high frequency generator
24 includes a burst generator 30 and burst multiplexer 32.
[0136] The burst multiplexer 32 receives burst parameters stored in
the control registers 18 for all the channels, selects the burst
parameters associated with an active channel, and outputs the
selected burst parameters to the burst generator 30. Specifically,
there is a pulse1/pulse2 burst register pair for each channel for
the burst option, totaling 32 registers for the 16 possible
channels in the embodiment shown. The burst multiplexer 32 is a
vector MUX that selects the pulse1/pulse2 register pair
corresponding to the active channel number. The select lines 34 to
the burst multiplexer 32 is the active channel number (ch_code)
from the arbitrator 22, which identifies the active channel at any
given time. The selected burst parameters for pulse1/pulse2 are
sent to the burst generator 30.
[0137] Within the burst generator 30, there are 2 independent burst
generator circuits, one for pulse1 and one for pulse2. The pulse1
and pulse2 signals belonging to the active channel are passed
through to the burst generator circuits from the arbitrator 22. If
the pulse1/pulse2 burst parameter data stored in the control
registers 18 is zero (therefore the selected burst parameters to
the burst generator 30 are also zero), then no burst is generated,
in which case the pulse1/pulse2 signals from the arbitrator 22 are
sent unaltered to the electrode driver 26. If the pulse1/pulse2
burst register in the control registers 18 is programmed, then the
pulse1/pulse2 duration will be replaced by (modulated to) the
corresponding programmed burst signal based on the selected burst
parameters from the burst multiplexer 32.
[0138] As an example, FIG. 27 illustrates that pulse1 for channel
one has been programmed for high frequency modulation while pulse2
for the same channel has not been programmed. Specifically, a
single pulse1 pulse has been modulated to (replaced with) five
higher frequency burst pulses. Thus, the frequency of the newly
modulated pulse1 is five times the frequency of the original pulse1
signal.
[0139] FIG. 23 is a more detailed functional block diagram of the
electrode driver 26 of FIG. 21. The electrode driver 26 includes a
pulse amplitude multiplexer 36, pulse scaler 38 and current drivers
40.
[0140] The pulse amplitude multiplexer 36 receives amplitude
parameters stored in the control registers 18 for all the channels,
selects the amplitude parameters associated with an active channel,
and outputs the selected amplitude parameters to the pulse scaler
38.
[0141] Specifically, 512 bytes (16 by 32 bytes--32 bytes for each
channel) in the control registers 18 are reserved for storing pulse
amplitude data. Each of the 16 channels is associated with 32 bytes
with each byte representing pulse amplitude information for pulse1
and pulse2 of each of the 32 electrodes E1-E32. From one byte, 7
bits are used to store the amplitude information for pulse1 and
pulse2 and the remaining bit (MSB) defines the polarity of the
pulse at the associated electrode as will be discussed later
herein.
[0142] Similar to the burst multiplexer 32, the pulse amplitude
multiplexer 36 is a vector MUX that selects the 32 bytes of
amplitude parameters for pulse1/pulse2 corresponding to the active
channel number. The select lines 34 to the amplitude multiplexer 36
is the active channel number (ch_code) from the arbitrator 22,
which identifies the active channel at any given time. The selected
amplitude parameters of pulse1/pulse2 for all 32 electrodes E1-E32
are sent to the pulse scaler 38. The pulse scaler 38 outputs
amplitude scaling factors for all electrodes of the active channel.
Thus, the pulse scaler 38 includes 32 identical scalers
corresponding to the 32 electrodes E1-E32. In the embodiment shown,
each scaler includes a D/A converter that converts the digital
amplitude value into a corresponding analog value I1-I32.
[0143] As discussed above, the signal generator 106 supports
asymmetrical pulse amplitude feature, which means the amplitude for
pulse1 and pulse2 can be different. The amplitude scaling data are
stored in the control registers 18. In the embodiment shown, 4 bits
are used to specify the scaling factor for pulse1 and pulse2 for
each electrode--2 bits for pulse1 and 2 bits for pulse2. Moreover,
the signal generator 106 can support asymmetrical pulse width
variation between pulse1 and pulse2.
[0144] The pulse scaler 38 adjusts the amplitude parameter by the
associated scaling factor stored in the associated 2 bits for
pulse1 and pulse2. In the embodiment shown, the pulse scaler 38
performs the scaling function by shifting to the right the content
of the selected amplitude parameter (7 bits of data from the
amplitude multiplexer 36) by the number stored in the corresponding
2 bit scaling factor. Since there are four possibilities in a 2 bit
number (0, 1, 2, or 3), the amplitude can be reduced by 1/2, 1/4 or
1/8.
[0145] For example, assume that the active channel is channel one,
the amplitude parameter for electrode E1 for channel one in the
control registers 18 is binary "1111111" (decimal 127) while the
associated 2 bit scaling factor is binary "11" for pulse1 and "01"
for pulse2. The number 127 represents 12.7 mA of current. For
pulse1, the 7 bit amplitude content will be shifted to the right by
3 bits (binary "11") for pulse1 and by 1 bit (binary "01") for
pulse2. Thus, the pulse scaler 38 will scale the amplitude of 127
down to 15 (binary "1111") for pulse1 and to 63 (binary "111111")
for pulse2, which corresponds to a current of 1.5 mA for pulse1 and
6.3 mA for pulse2.
[0146] In the embodiment shown, the scaling factor parameters are
stored in unused bits of burst parameters and are passed to the
pulse scaler 38 from the burst multiplexer 32. However, dedicated
memory can be allocated in the control registers 18. It is to be
noted that the stimulation programming software should ensure that
at any given instant of time, the algebraic sum of all electrode
currents is zero, and the dc average of the current per cycle at
each electrode is also zero.
[0147] Another function performed by the pulse scaler 38 is that it
converts the scaled amplitude digital value into an analog output
current that is linearly proportional to the value of the digital
value. The analog output current for each electrode is supplied to
the current drivers 40. In the embodiment shown, the analog output
current represents 1/20 of the actual current to be supplied to the
associated electrode.
[0148] The current drivers 40 amplify the analog output currents
from the pulse scaler 38 and switches the amplified current to the
appropriate electrodes E1-E32 based on the pulse parameters stored
in the control registers 18. Since there are 32 electrodes, there
are 32 electrode drivers 40 in the embodiment shown. In the
embodiment shown, each current driver 40 is a current driver that
amplifies the input signal from the pulse scaler 38 by 20
times.
[0149] FIG. 24 is a functional illustration of two of the current
drivers of FIG. 23. Each electrode driver 40 can be a current
source that can sink or source current whose amplitude is based on
a control current signal coming from the pulse scaler 38. In the
embodiment shown, the current source includes a pair of NMOS
current source 42 and PMOS current source 44 that are coupled in
series between the voltage supply and ground. Each of the PMOS and
NMOS current sources can be implemented as a current mirror in a
well-known manner. The NMOS current source 44 sinks current from
the associated electrode to ground while the PMOS current source 44
sources current from the positive voltage supply to the associated
electrode. Each electrode driver 40 includes switches 46 and 48
connected in series between the voltage supply and ground to either
source or sink the current.
[0150] The control registers 18 store the pulse parameter that
relate to whether a particular electrode will be sourcing current
or sinking current. In the embodiment shown, bit 7 (MSB) of each
byte of the pulse amplitude parameters that are supplied to the
burst multiplexer 32 is used to specify whether a particular
current source 40 will be sourcing current or sinking current. If
the bit is zero, during pulse1, switch 46 will be turned on while
switch 48 will be turned off, and during pulse2, switch 46 will be
turned off while switch 48 will be turned on. If the bit is set
(i.e., it is a "1"), during pulse1, switch 46 will be turned off
while switch 48 will be turned on, and during pulse2, switch 46
will be turned on while switch 48 will be turned off.
[0151] FIG. 24 illustrates the current path for two electrodes Ej
and Ek when bit 7 for Ej=0 and bit 7 for Ek=1. During pulse1,
switch 46 for Ej and switch 48 for Ek turn on to create a current
path 50. In that instance, the PMOS current source 42 for Ej will
be sourcing current while the NMOS current source 44 for Ek will be
sinking current to ground. Conversely, during pulse2, switch 46 for
Ek and switch 48 for Ej turn on to create a current path 52. In
that instance, the PMOS current source 42 for Ek will be sourcing
current while the NMOS current source 44 for Ej will be sinking
current to ground.
[0152] The pulse shape at the electrode Ej will look similar to the
E1 waveform as shown in FIG. 26 while the pulse shape at the
electrode Ek will look similar to the E2 waveform.
[0153] The foregoing specific embodiments represent just some of
the ways of practicing the present invention. Many other
embodiments are possible within the spirit of the invention.
Accordingly, the scope of the invention is not limited to the
foregoing specification, but instead is given by the appended
claims along with their full range of equivalents.
* * * * *